An “endoprosthesis” corresponds to an artificial device that is placed inside the body, more particularly, within an anatomical lumen. A “lumen” refers to a cavity of a tubular organ such as a blood vessel. A stent is an example of an endoprosthesis. Stents are generally cylindrically shaped devices, which function to hold open and sometimes expand a segment of a blood vessel or other anatomical lumen such as urinary tracts and bile ducts. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels.
The treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent. “Delivery” refers to introducing and transporting the stent through an anatomical lumen to a region, such as a lesion, in a vessel that requires treatment. “Deployment” corresponds to the expanding of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into an anatomical lumen, advancing the catheter in the anatomical lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen.
In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter. Mounting the stent typically involves radially compressing or crimping the stent onto the balloon. The stent is then expanded by inflating the balloon. The balloon may then be deflated and the catheter withdrawn.
The stent must be able to satisfy a number of mechanical requirements. First, the stent must be capable of withstanding the structural loads, namely radial compressive forces, imposed on the stent as it supports the walls of an anatomical lumen. Therefore, a stent must possess adequate radial strength. Radial strength, which is the ability of a stent to resist radial compressive forces, is due to strength and rigidity around a circumferential direction of the stent. Radial strength and rigidity, therefore, may also be described as hoop strength and rigidity.
Once expanded, the stent must adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. For example, a radially directed force after deployment may cause a stent to plastically deform, which can reduce clinical effectiveness.
In addition, the stent must possess sufficient flexibility to allow for crimping, deployment, and cyclic loading after deployment. Longitudinal flexibility is important to allow the stent to be maneuvered through a tortuous anatomical path and to enable it to conform to a deployment site that may not be linear or may be subject to flexure. Also, the stent must be biocompatible so as not to trigger any adverse responses.
The structure of a stent typically comprises scaffolding that includes a pattern or network of interconnecting structural elements often referred to in the art as struts, links and rings. The scaffolding is designed so that the stent can be radially compressed (to allow crimping) and radially expanded (to allow deployment).
Polymers have been used to make stent scaffolding. The art recognizes a variety of factors that affect a polymeric stent's ability to retain its structural integrity when subjected to external loadings, such as crimping and balloon expansion forces. These interactions are complex and the mechanisms of action not fully understood. According to the art, characteristics differentiating a polymeric, bio-absorbable stent scaffolding of the type expanded to a deployed state by plastic deformation from a similarly functioning metal stent are many and significant. Indeed, several of the accepted analytic or empirical methods/models used to predict the behavior of metallic stents tend to be unreliable, if not inappropriate, as methods/models for reliably and consistently predicting the highly non-linear behavior of a polymeric load-bearing, or scaffolding portion of a balloon-expandable stent. The models are not generally capable of providing an acceptable degree of certainty required for purposes of implanting the stent within a body, or predicting/anticipating the empirical data.
Polymer material considered for use as a polymeric stent scaffolding, such as Poly L-lactic acid (PLLA) and poly lactic-co-glycolic acid (PLGA), may be described through comparison with a metallic material conventionally used to form stent scaffolding. In comparison to metals, a suitable polymer has a low strength to weight ratio, which means more material is needed to provide an equivalent mechanical property to that of a metal. Therefore, struts in polymeric scaffolding must be made thicker and wider to have the strength needed. Polymeric scaffolding also tends to be brittle or have limited fracture toughness. The anisotropic and rate-dependant inelastic properties (i.e., strength/stiffness of the material varies depending upon the rate at which the material is deformed) that are inherent in the material only compound this complexity in working with a polymer, particularly, a bio-absorbable polymer such as PLLA and PLGA.
Therefore, processing steps performed on and design changes made to a metal stent that have not typically raised concerns for unanticipated changes in the average mechanical properties, may not also apply to a polymer stent due to the non-linear and sometimes unpredictable nature of the mechanical properties of the polymer under a similar loading condition. It is sometimes the case that one needs to undertake extensive validation before it even becomes possible to predict more generally whether a particular condition is due to one factor or another—e.g., was a defect the result of one or more steps of a fabrication process, or one or more steps in a process that takes place after stent fabrication, e.g., crimping. As a consequence, a change to a fabrication process, post-fabrication process or even relatively minor changes to a stent pattern design must, generally speaking, be investigated more thoroughly than if a metallic material were used instead of the polymer. It follows, therefore, that when choosing among different polymeric stent designs for improvement thereof, there are far less inferences, theories, or systematic methods of discovery available, as a tool for steering one clear of unproductive paths, and towards more productive paths for improvement, than when making design changes in a metal stent.
It is recognized, therefore, that, whereas inferences previously accepted in the art for stent validation or feasibility when an isotropic and ductile metallic material was used, such inferences would be inappropriate for a polymeric stent. A change in a polymeric stent pattern may affect, not only the stiffness or lumen coverage of the stent in its deployed state, but also the propensity for fractures to develop when the stent is crimped or being deployed. This means that, in comparison to a metallic stent, there is generally no assumption that can be made as to whether a changed stent pattern may not produce an adverse outcome, or require a significant change in a processing step (e.g., tube forming, laser cutting, crimping, etc.). Simply put, the highly favorable, inherent properties of a metal (generally invariant stress/strain properties with respect to the rate of deformation or the direction of loading, and the material's ductile nature), which simplify the stent fabrication process, allow for inferences to be more easily drawn between a changed stent pattern and/or a processing step and the ability for the metallic stent to be reliably manufactured with the new pattern and without defects when implanted within a living being.
A change in the pattern of the struts and rings of a polymeric stent scaffolding that is plastically deformed, both when crimped to, and when later deployed by a balloon, unfortunately, is not as easy to predict as a metal stent. Indeed, it is recognized that unexpected problems may arise in polymer stent fabrication steps as a result of a changed pattern that would not have necessitated any changes if the pattern was instead formed from a metal tube. In contrast to changes in a metallic stent pattern, a change in polymer stent pattern may necessitate other modifications in fabrication steps or post-fabrication processing, such as crimping and sterilization.
A problem encountered with polymeric stents after they are crimped onto a balloon is the development of fractures and other defects that require the stent to be rejected and scrapped. Cracks and other defects can render the stent incapable of functioning properly when fully deployed by the balloon. Another problem is that deployment of polymeric stents from the crimped state to a deployed state in a patient can produce strain that adversely affects the ability of the stent to stay in the deployed state and remain at the implantation site, especially under cyclic loading conditions inherent in a patient's circulatory system. The strain induced during deployment can result in significant loss in radial strength.
In light of the foregoing, there is a need for a stent pattern and manufacturing method that reduces the cracks and other defects due to crimping. There is also a need for a stent pattern and manufacturing method that results in less strain when a stent is deployed for implantation.